Dental Laser System and Treatment Method

ABSTRACT

An improved dental laser system has been developed to cut enamel quickly and precisely, without detrimental residual energy, to provide a replacement for conventional high speed rotary burrs and commercially available dental laser systems.

CROSS-REFERENCES TO RELATED APPLICATIONS

This application is a continuation-in-part of and claims priority under35 U.S.C. §120 to co-pending, commonly owned U.S. patent applicationSer. No. 12/847,739, entitled DENTAL LASER SYSTEM USING MIDRANGE GASPRESSURE, filed Jul. 30, 2010, which claims priority to and the benefitof U.S. provisional patent application Ser. No. 61/229,997, filed Jul.30, 2009. The entire disclosures of these two applications are herebyincorporated herein by reference in their entireties.

FIELD OF INVENTION

The present invention relates to systems and methods for removing decay,cutting, drilling or shaping hard tissue, removing and cutting softtissue, modifying hard tissue for caries inhibition and modifying hardtissue surface conditions to aid in adhesion to hard tissue. The presentinvention applies to oral tissue, gums and teeth, e.g., to human oranimal oral tissue, gums and teeth, and other biological materials.

BACKGROUND OF THE INVENTION

A tooth has three layers. The outermost layer is the enamel which is thehardest and forms a protective layer for the rest of the tooth. Themiddle and bulk of the tooth is made up of the dentin, and the innermostlayer is the pulp. The enamel and dentin are similar in composition andare roughly 85% mineral, carbonated hydroxyapatite, while the pulpcontains vessels and nerves which are sensitive to pressure andtemperature. In this application of drilling or contouring orconditioning the enamel and dentin, the pulp's temperature sensitivityis of concern. A rise in temperature of 5.5° Celsius can lead topermanent damage of the tooth's pulp.

Over the last 10 to 15 years, research has taken place to define laserparameters that allow the enamel and dentin of a tooth to be removed,drilled, contoured or conditioned, all being removal processes, withoutheating the pulp. Ideally the laser pulses should vaporize the enameland dentin converting the mass to gas with minimal residual energyremaining in the dentin to heat the pulp.

The use of lasers in dentistry has been considered since theintroduction of the laser. Dental lasers used to drill and cut were theinitial applications. High energy density pulses were initially used,but these pulses could potentially damage the tooth pulp or soft tissue,so lower energy pulse configurations were explored. With lower peakpower energy pulses longer pulse times were used, which affected thetooth enamel detrimentally.

Various laser wavelength interactions were explored, UV to the FarInfrared, to understand the optical coupling efficiencies. Opticalcoupling was found to be critical with the greatest coupling being inthe 2.7-3.0 μmeter and 9.3-9.6 μm wavelength ranges. When reflectance isconsidered, the 9.3-9.6 μmeter range was found to couple up to 3 timesbetter than any other wavelength range.

Having identified the most effective coupling wavelength, the time andthreshold to ablate hard tissue had to be determined. Research has shownthat the thermal relaxation time of hard tissue is 1 to 2 μsec with athreshold ablation energy of approximately 5 mJ (milli-Joules).

Recognizing the need for laser pulses in the 9.3 to 9.6 μmeterwavelength range with microsecond pulse widths and pulse energies of 5to 15 mJ, DC excited TEA (transversely excited atmospheric) lasers wereadopted. Since the TEA lasers have a very short pulse length, i.e.,hundreds of nanoseconds, the TEA lasers were modified for long pulseoperation and modified pulse shapes. Additionally a RF (Radio Frequency)CW (continuous wave) laser was studied, but its shortest pulse lengthwas only 50 μseconds, so the pulses heated the hard tissue significantlymore than the shorter pulse widths.

To date, RF excited CO₂ CW lasers seeking the greatest RF to Opticalefficiency typically operate at 70 to 100 Torr (or about 9,332-13,332Pascals (Pa)) and the shortest pulse lengths produced are typically 50μseconds. Typical gas pressure for a normal RF excited CO₂ laser, usedin the prior art, is 80 Torr (or about 10,665 Pa). CO₂ TEA lasersoperating at atmospheric pressure produce 9.3 to 9.6 μmeter pulses athundreds of nanoseconds in pulse length. TEA lasers generally do notoperate in sealed operation, do not have long operating lifetimes orhigh pulse repetition rates, and are expensive to package. While a “longpulse” TEA laser can be manufactured to produce the optimal CO₂ laserpulsing parameters, TEA lasers are larger and more expensive than RFexcited lasers and therefore are not an ideal match for a dental laserapplication—where size and cost are critical. None of the approaches todate, therefore, have produced a full set of optimal parameters in acommercially acceptable format for effectively working with enamel anddentin, without heating the pulp.

SUMMARY OF THE INVENTION

In accordance with one aspect a method for treating hard tissue isprovided, the method including generating a laser beam with a repetitionrate greater than about 0.5 kHz, directing the laser beam at an exposedsurface of the hard tissue (e.g., tooth enamel, dentin, etc.), andremoving at least a portion of the exposed surface of the hard tissuewithout substantially increasing temperature of adjacent tissue (e.g.,tooth pulp). The laser beam may have one or more of: a wavelength in arange of about 9 μm to about 10 μm, a pulse length in a range of up toabout 30 μsec, a fluence up to about 50 J/cm², a repetition rate up toabout 4 kHz, and a nominal diameter of up to about 2000 μm. In someembodiments, the laser beam of the method for treating hard tissueincludes laser pulses such that a pulse energy associated with eachpulse is in a range of up to about 30 mJ.

In some embodiments, the laser beam of the method for treating hardtissue is generated by a CO₂ laser. The CO₂ laser may operate with a gasin a range of about 260 Torr to about 600 Torr, and the gas may be¹²C(¹⁸O)₂ gas or ¹²C(¹⁶O)₂ gas.

In some embodiments, the laser beam is directed at the exposed surfaceof the hard tissue utilizing a hollow waveguide. The positioning of thehollow waveguide relative to the exposed surface of the hard tissue maybe controlled manually or robotically. The laser beam may be movedrelative to the exposed surface of the hard tissue to removeincrementally the portion of the exposed surface of the hard tissue. Insome embodiments, a remaining portion of the hard tissue does notexhibit any substantial charring. The hard tissue is removed at a rateof up to about 0.5 g/sec, or at a volumetric rate of up to about 1.7cm³/sec. The hard tissue may be removed by explosive vaporization.

In some embodiments, the method for treating hard tissue includesdirecting a fluidic flow at the exposed surface of the hard tissue. Thefluidic flow may be selected from the group consisting of a fluid, agas, and combinations thereof. The fluidic flow may be adapted tocontribute to either controlling temperature of the hard tissue, orreducing redeposition of removed material onto the exposed surface ofthe hard tissue, or both. The fluidic flow may be directed to theexposed surface of the hard tissue at an angle of incidence in a rangefrom about zero degrees to about 90 degrees. The hard tissue treated bythe method may be dental enamel, dentin, bone, or combination thereof.

In some embodiments, the exposed surface of the hard tissue at which thelaser beam is directed during a single pulse has a cross-sectional areaof up to about 0.03 cm². In some embodiments, directing the laser beamincludes delivering a sequences of pulses for up to about three minutes.The laser beam may be directed to the exposed surface of the hard tissueat an angle of incidence up to about 90 degrees. Directing the laserbeam may include focusing the laser beam using a lens having a focallength, and positioning the exposed surface of the hard tissue relativeto the lens at a distance slightly less than the focal length. In someembodiments, directing the laser beam includes modifying the laser beamusing an optical system, such that a profile of the modified laser beamis non Gaussian.

In accordance with another aspect, a system for treating hard tissue isprovided, the system including a laser beam generator for generating alaser beam with a repetition rate greater than about 0.5 kHz. The systemalso includes an optical component for directing the laser beam at anexposed surface of the hard tissue, and a laser beam controller forcontrolling at least one laser beam parameter, such that at least aportion of the exposed surface of the hard tissue is removed withoutsubstantially increasing a temperature of adjacent tissue.

The laser beam parameters may be one or more of: a wavelength in a rangeof about 9 μm to about 10 μm, a pulse length in a range of up to about30 μsec, a fluence up to about 50 J/cm², a repetition rate up to about 4kHz, a nominal diameter of up to about 2000 μm, and a pulse energy in arange of up to about 30 mJ.

The laser beam generator of the system for treating hard tissue may be aCO₂ laser beam generator. The CO₂ laser beam generator may operate witha gas in a range of about 260 Torr to about 600 Torr, and the gas may be¹²C(¹⁸O)₂ gas or ¹²C(¹⁶O)₂ gas.

In some embodiments, the optical component of the system for treatinghard tissue includes a hollow waveguide. The system for treating hardtissue may include a handpiece for manually positioning the hollowwaveguide relative to the exposed surface of the hard tissue. In someembodiments, the system includes a robotic controller for positioningthe hollow waveguide relative to the exposed surface of the hard tissue.The robotic controller may be configured to move the laser beam relativeto the exposed surface of the hard tissue to remove incrementally theportion of the exposed surface of the hard tissue.

In some embodiments, the hard tissue is removed at a rate of up to about0.5 g/sec, or at a volumetric rate of up to about 0.17 cm³/sec. The hardtissue may be removed by explosive vaporization. A remaining portion ofthe hard tissue may not exhibit any substantial charring.

In some embodiments, the system for treating hard tissue includes afluid dispenser for directing a fluidic flow at the exposed surface ofthe hard tissue. The fluidic flow dispensed by the fluid dispenser maybe a fluid, a gas, or combinations thereof. The fluidic flow may beadapted to contribute to either controlling temperature of the hardtissue, or reducing redeposition of removed material onto the exposedsurface of the hard tissue, or both. The fluid dispenser may direct thefluidic flow to the exposed surface of the hard tissue at an angle ofincidence in a range from about zero degrees to about 90 degrees. Thehard tissue may be dental enamel, dentin, bone, or combination thereof.

In some embodiments, the laser beam controller of the system fortreating hard tissue configures the laser beam as a sequences of pulsesfor up to about three minutes. The exposed surface of the hard tissue atwhich the laser beam may be directed during a single pulse may have across-sectional area of up to about 0.03 cm².

In some embodiments, the optical component is configured to direct thelaser beam to the exposed surface of the hard tissue at an angle ofincidence up to about 90 degrees. The optical component may include alens having a focal length for focusing the laser beam, such that theexposed surface of the hard tissue relative to the lens is at a distanceslightly less than the focal length. In some embodiments, the opticalcomponent modifies the laser beam, such that a profile of the modifiedlaser beam is non Gaussian.

DESCRIPTION OF THE DRAWINGS

The present invention will become more apparent in view of the attacheddrawings and accompanying detailed description. The embodiments depictedtherein are provided by way of example, not by way of limitation,wherein like reference numerals refer to the same or similar elements.The drawings are not necessarily to scale, emphasis instead being placedupon illustrating aspects of the invention. In the drawings:

FIG. 1 is a block diagram of an embodiment of a dental laser system, inaccordance with aspects of the present invention.

FIG. 2 is a flowchart of an embodiment of outputting laser opticalenergy from a CO₂ dental laser system, in accordance with aspects of thepresent invention.

FIG. 3 is a table of typical laser parameters and dental systemperformance parameters in accordance with the present invention.

FIG. 4 depicts absorption of radiation by tooth enamel at variouswavelengths and pulse widths.

FIG. 5 is a table of the laser parameters and performance parametersmeasured during the operation of an experimental laser system inaccordance with the present invention.

FIG. 6 depicts the rates of removal of tooth enamel by weightcorresponding to treatment using laser beams having different parametersin accordance with the present invention.

FIG. 7 presents a table of the rates of removal of tooth enamel ondifferent molars corresponding to one set of laser parameters inaccordance with the present invention.

FIG. 8 shows the temperatures observed by various thermocouples duringtreatment using the experimental system in accordance with the presentinvention.

FIG. 9 depicts relative performances of an exemplary laser systemaccording to the present invention, and conventional systems.

DETAILED DESCRIPTION

Hereinafter, aspects of the present invention will be described byexplaining illustrative embodiments in accordance therewith, withreference to the attached drawings. While describing these embodiments,detailed descriptions of well-known items, functions, or configurationsare typically omitted for conciseness.

It will be understood that when an element is referred to as being “on”or “connected” or “coupled” to another element, it can be directly on orconnected or coupled to the other element or intervening elements can bepresent. In contrast, when an element is referred to as being “directlyon” or “directly connected” or “directly coupled” to another element,there are no intervening elements present. Other words used to describethe relationship between elements should be interpreted in a likefashion (e.g., “between” versus “directly between,” “adjacent” versus“directly adjacent,” etc.).

The terminology used herein is for the purpose of describing particularembodiments only and is not intended to be limiting of the invention. Asused herein, the singular forms “a,” “an” and “the” are intended toinclude the plural forms as well, unless the context clearly indicatesotherwise. It will be further understood that the terms “comprises,”“comprising,” “includes” and/or “including,” when used herein, specifythe presence of stated features, steps, operations, elements, and/orcomponents, but do not preclude the presence or addition of one or moreother features, steps, operations, elements, components, and/or groupsthereof.

With respect to dental laser systems, the wavelength with the highestabsorption in hydroxyapatite has been determined to be in the 9.3 to 9.6μmeter range and the thermal relaxation time of hydroxyapatite to be amaximum of 2 μseconds at the 9.3 to 9.6 μm wavelength range. Therefore,the ideal pulse parameters for removing the hydroxyapatite appear to be9.3 to 9.6 μmeter energy in a less than 50 μsecond format. In accordancewith one preferred embodiment, a laser is provided that produces a beamhaving pulse parameters for removing hydroxyapatite using 9.3 to 9.6 μmwavelength energy in a less than 50 μsecond format.

The 9.3 to 9.6 μm energy is typically produced using a CO₂ laser with alaser gas mixture of ¹²C(¹⁸O)₂, wavelength selective resonator optics,more expensive inter-cavity wavelength devices, or a combination of thethree. In accordance with the present invention, the 50 μsecond pulsesare produced with a fast pulse rise and fall time, which is effected bylaser gas pressure. In order to produce pulses of less than or equal to50 μseconds in length, gas pressure of at least about 260 Torr (or about34,663 Pa) is used.

According to one preferred embodiment, a CO₂ gas laser, in either awaveguide or slab resonator format, filled with gas that is in a rangeof about 260 Torr to about 600 Torr (or about 34,700-80,000 Pa), is RFexcited for use in all dental applications. A range of about 260 toabout 600 Torr (or about 34,700-80,000 Pa) may be preferable in manydental applications. Since waveguide and slab resonators are generallyknown in the art, they are not discussed in detail herein.

In some embodiments, the pressure can be in a range of about 280-550Torr (or about 37,330-73,327 Pa), or about 300-500 Torr (or about39,996-66,661 Pa), about 320-450 Torr (or about 42,663-59,995 Pa), about340-400 Torr (or about 45,329-53,328 Pa), as examples.

The laser can be operated in CW or pulsed mode for cutting and drillingapplications, respectively. DC and RF power supplies are configured toaid in low power CW operation, while supporting high peak power pulseoperation. The laser output is coupled to a beam delivery system todeliver the optical energy to the patient. The laser provides the 9.3 to9.6 nm energy wavelength, with a fast pulse rise and fall time (e.g.,not more than about 50 μseconds, and typically not more than 20μsecond), sealed off operation, high repetition rates in a smallreliable package.

FIG. 1 shows an embodiment of a dental laser system 100 according toaspects of the present invention. In the embodiment of FIG. 1, a DCpower supply 10 is provided that rectifies as AC input power (notshown). In one preferred embodiment, the DC power supply 10 is comprisedof a continuous wave (CW) DC section 12 and a pulsed DC section 14. TheDC section 12 is sized to run the laser for CW applications, such assoft tissue cutting, and the peak power DC section 14 supplies the peakenergy for pulsing applications, such as hard tissue or bonemodification.

Item 20 is a radio frequency (RF) power supply for the conversion of theDC energy to RF energy in the 40 to 125 MHz range. Item 30 is a CO₂laser with the RF energy as an input and 9.3 to 9.6 μmeter opticalenergy as an output, via an output coupler 32. And item 40 is a beamdelivery apparatus, which delivers the optical energy from the laser toitem 50, which represents a patient's mouth.

CO₂ laser 30 in this embodiment includes a rear mirror 34 and a laserdischarge area 36. The mirror 34 directs optical energy through thelaser discharge area 36. The output coupler 32 couples the beam out ofthe laser. In this case the laser is a gas laser, so the output couplercouples the beam out of the laser without allowing the laser gas out.The CO₂ laser 30 also includes a laser gas pressure vessel 38 that isfilled with a gas at a pressure in a range of about 260 to about 600Torr (or about 34,700-80,000 Pa).

The output laser energy is provided to the beam delivery apparatus 40,where it can then be directed to a target, such as a patient's mouth. Inthis embodiment, the beam delivery apparatus 40 can include flat orcurved mirrors or a combination configured to steer optical energyoutput from the CO₂ laser.

In this exemplary configuration, the dental laser system 100 can operateat both low power CW operation, e.g., <10 watts, for the cutting of gumsand oral tissue, and high peak power pulsing operation, e.g., >5 mJpulse energy at 1 to 50 μseconds pulse widths up to 10 kHz. The CO₂laser 30 can operate at wavelengths between 9 and 11 μm. The lasersystem 100 preferably provides high peak power pulsing operation at theideal absorption wavelength for the hydroxyapatite in dental hardtissues. The pulse widths and pulse energy are ideally suited to ablatehydroxyapatite, leaving very little residual heat in the tooth to damagethe pulp even up to 10 kHz in operation.

FIG. 2 is an embodiment of a method of outputting laser optical energyfrom a CO₂ dental laser system. The method 200 includes providing adirect current (DC) power supply in step 210, providing a radiofrequency (RF) power supply coupled to the DC power supply in step 220,filling a CO₂ laser with gas at a pressure in a predetermined pressurerange (e.g., about 260 to about 600 Torr (or about 34,700-80,000 Pa)) instep 230, and steering the laser optical energy output from the from theCO₂ laser to a patient using a beam delivery system 240.

FIG. 3 is a table listing a number of typical laser operationalparameters and associated dental system performance parameters discussedin more detail, below. Values are provided for minimum and maximumconditions, as well as nominal conditions contemplated for use withvarious embodiments and in various applications of the invention. Itshould be noted that the minimum and maximum values are not boundariesand that actual values may be lower or higher by 10% or 20% or more ofthe total range for each parameter.

More specifically, a hard tissue, such as tooth enamel and dentin,comprises water and minerals. While extensive dental laser studies havebeen conducted at hard tissue's water and mineral absorption bands, theresearch emphasis has been on the efficiency of absorption. Dentalpractices are primarily concerned with cutting speed and resolution.Building on the efficiency results from prior research, an exemplarydental laser system was built with the flexibility to change pulse widthand average power levels at hard tissue's mineral absorption band toimprove cutting efficacy, as described in detail below in the Example.

Cutting or drilling hard tissue in a tooth primarily equates to removingthe tissue's hardest outer layer, the enamel. Enamel is a biologicalcomposite generally containing by volume approximately 12% water,approximately 85% mineral (carbonated hydroxyapatite), and about 3%protein and lipids. Accordingly high efficacy may be achieved by cuttingor drilling the 85% mineral constituent of the enamel. It has been shownthat dental enamel's mineral content has a peak absorption (8000 cm⁻¹)at 9.6 μm, absorbing electromagnetic radiation up to ten times higherthan at 10.6 μm or other laser wavelengths in the visible or other IRregion. In the exemplary system described in the Example, a 9.4 μm laserwas used as the “proof of concept” because at present it is acommercially available laser having a wavelength close to 9.6 μm. FIG. 4shows that the difference in absorption at 9.4 μm and 9.6 μm is notsignificant.

It is desirable that the laser parameters for dental hard tissueablation, namely pulse duration and laser wavelength, be selected suchthat practical ablation rates can be achieved while minimizing theresidual energy, ER, deposition in the tooth. Residual energy can be theenergy directed by a laser beam to the hard tissue (e.g., tooth enamel)that does not cause tissue ablation, but instead, is absorbed into thesurrounding tissue. In general, the absorption of radiation in enamel isheterogeneous due to the inherent microstructure of the enamel. As aresult, the mechanism of ablation varies with the nature of the primaryabsorber in which the laser radiation is absorbed. For example,absorption in water results in water mediated ablation and absorption inthe bulk of the enamel rods leads to melting and explosive vaporization.As described above, the absorption at 9.6 μm occurs in the enamel rodmineral content typically causing explosive vaporization, as explainedbelow. Because the tooth enamel includes approximately 85% mineral,ablation efficiency of enamel is usually high at or near the mineralabsorption wavelength of 9.6 μm.

Due to the strong wavelength absorption of CO₂ laser radiation, thelaser energy is usually deposited in a very thin absorption layer at thesurface of the hard tissue (e.g., tooth enamel). This can result in astrong internal pressure build-up in a small volume of the hard tissue.If the pressure rise time is shorter than the thermal relaxation time ofthe tissue (e.g., the time during which the heat generated in the hardtissue is dissipated), the pressure finally exceeds the ultimate tensilestrength of the tissue causing the tissue to tear apart in a localizedmicro-explosion, also called explosive vaporization. The excess heat isejected together with heated ablation debris and is thus removed fromthe tissue. Accordingly, in explosive vaporization, thermal damage tothe surrounding tissue can be prevented. CO₂ lasers may be suitable forexplosive vaporization, because their wavelength is strongly absorbed inthe mineral component (hydroxyapatite) of bone tissue, such that a smalltissue volume can be heated extremely quickly.

As described above, explosive vaporization may occur when the pressurerise time is shorter than the thermal relaxation time of the tissue. Dueto the nature of explosive vaporization, the incident laser energy iseither consumed and ejected in the ablation process, or absorbed in thetissue in the form of heat, or residual energy ER. The residual energycan cause heating of the surrounding tissue (e.g., internal tissue ofthe tooth) as opposed to the removal of hard tissue. Therefore,minimizing the ER can increase ablation efficiency.

A desired pulse duration or pulse length can be on the order of therelaxation time for axial heat conduction (τz) of the deposited energyin the tissue surface. For laser pulse durations substantially longerthan τz, the laser energy may be conducted away from the enamel surfaceinto the interior pulp of the tooth during the laser pulse, resulting ininefficient surface heating. If the pulse duration is too short,however, the required power density to cause explosive vaporization maybe too high, likely causing the generation of a plasma in the plume ofablated material that shields the surface and reduces the efficiency ofablation.

The thermal relaxation time of the deposited laser energy in enamel atthe wavelength of approximately 9.6 μm is about 1 μsec. Thereforeablation efficiency of hard tissue using a radiation wavelength at ornear the mineral absorption wavelength of 9.6 μm, may be increased usingpulses approximately 1 μsec in duration. In one experiment the lowestresidual energy was noted at 9.6 μm wavelength, using 5 μsec pulselength and 15 J/cm² fluence (described below) where less than 25% of theincident energy was absorbed as residual energy.

Though pulses of short duration can increase the ablation efficiency bydecreasing the amount of residual energy absorbed, the longer the pulselength the lower the cost per watt of average laser power. Therefore, itis beneficial to select a pulse length that can simultaneously decreasethe amount of residual energy absorbed and the cost per watt of thelaser power. Accordingly, it is important to understand the residualenergy, ER, as a function of pulse length. The shorter laser pulses,e.g., 5-20 μsecs, ablate at a significantly lower ablation threshold andleave a lower residual energy than the longer laser pulses that canachieve higher ablation rates per pulse. Pulse lengths of less than orequal to 20 μsecs with fluences of 10 J/cm² or greater lead to lowresidual energy and high cutting efficiency. Longer pulse lengths, e.g.,longer than 20 μsec, may lead to higher ablation rates per pulse, buthigher residual energy and lower ablation efficiency. Shorter pulselengths, e.g., 5 to 20 μsec in duration, may have higher ablationefficiency and lower residual energy, but may result in lower ablationrates per pulse. Duty cycle (i.e., pulse length divided by total timebetween pulses can be varied as desired. See FIG. 3 for typical dutycycle values.

In addition to the absorption wavelength and the pulse length requiredfor high ablation efficiency, the hard tissue enamel rod heating speed,leading to explosive vaporization, generally depends on pulse energydensity, commonly referred to as fluence. In other words, the totalenergy delivered to a target site by one laser pulse may be expressed asthe product of fluence and area of the treatment site. A threshold,i.e., minimal radiation exposure φ_(th) (J/cm²) is theoreticallynecessary to induce an ablating micro-explosion at a treatment site bycausing a sufficient pressure build-up. Taking into account the lossesdue to diffusion or reflection of radiation, however, the actual fluencerequired to be directed toward the enamel may be greater than thetheoretical minimum. For example, to accumulate the necessary energy inthe absorption layer of the hard tissue to reach the theoretical energythreshold, the energy may need to accumulate fast enough so as tocompensate losses from thermal diffusion. Accordingly, there is aminimum required fluence threshold. For example, the ablation thresholdsare about 0.5 and 3 J/cm² for 9.6 μm and 10.6 μm, respectively, aftercorrection for reflectance losses. A desirable fluence is above thethreshold for explosive vaporization, but one that minimizes theresidual energy, ER, leading to high ablation efficiency. Consideringthe 1 μsec thermal relaxation time and the 0.5 J/cm² ablation threshold,a low residual energy in one embodiment occurs at 5 μsec pulse lengthand 15 J/cm² fluence. The ablation rate of enamel can saturate above 25J/cm², and, hence a desired fluence is generally between about 1 andabout 25 J/cm². Note, however, that higher fluences, up to about 50J/cm² can be used. Fluences lower than 1 J/cm² can be used, for example,to pit the surface of the enamel to increase bonding of a cover layer orcoating.

Generally, the speed of ablation of hard tissue depends on the averagelaser power used. Average laser power can be expressed as the product ofpower per pulse and the number of pulses, or power per pulse times thepulse repetition rate. Longer pulses, e.g. longer than 25 μsec induration, can ablate more material per pulse, but shorter pulses, e.g.,shorter than 20 μsec in duration, are more efficient at ablating dentalhard tissue. Assuming similar average powers, ablation rates can bedetermined by the product of the ablation per pulse and the pulserepetition rate. Therefore shorter pulse lengths having lower residualenergy levels can lead to higher dental cutting efficacy because morepulses per second, a higher repetition rate, can be utilized for thesame average power. Without practical concerns, higher average powerslead to faster dental laser system ablation rates. But, higher laseraverage power may lead to a higher product cost, and the human handlimits scanning efficiency above repetition rates of about 1 to 2 kHz.Rates up to 4 kHz may be employed. Additionally, computer controlledrobotic systems may be employed advantageously, to precisely control theapplication and location of the laser beam to the hard tissue, e.g., tocut a contour to receive a crown or pre-formed filling.

A 9.3 μm wavelength experimental dental system used pulse widths of 35to 75 μsec with fluences of 1 to 6 J/cm², and repetition rates up to 400Hertz. This dental laser system, with an integrated scanner, waseffective in cutting dental hard tissue with substantially notemperature rise, but cut very slowly. The long pulse lengths and lowfluence led to cutting speeds that are not desirable for generalclinical use, and are not significantly better than the conventionaldrilling systems using rotary dental burrs.

Many known dental laser systems operate by exploding the water in hardtissue at wavelengths of 2.7 to 3.0 μm. Cutting hard tissue throughwater absorption is referred to as “thermo-mechanical” or“water-mediated” ablation. During rapid heating, the inertially confinedwater can create enormous substructure pressures that can lead to theexplosive removal of the surrounding mineral matrix. Several studies ofhard tissue ablation in the radiation wavelength λ at or near 3.0 μmindicate that large particles are ejected with high velocity, whichstrongly supports the mechanism of a water-mediated explosion process.One commercially available 2.78 μm wavelength dental laser systemoperates at relatively long pulse widths of 140 μsec, fluence of 32 to60 J/cm², having a repetition rate of up to 50 Hertz. The relativelyhigh pulse energy of this system requires a low repetition ratecorresponding to an equivalent average power level. The longer pulselengths, and associated higher pulse energies, also lead to higherresidual energy, ER, in the tooth.

EXAMPLE

An experimental dental laser system was constructed to operate at 9.4 μmwavelength with pulse widths of about 5 to 20 μsec, fluence of about 4to 15 J/cm², and repetition rates of about 250 to 2000 Hertz. Theparameters of this dental laser system, such as wavelength, pulse width,fluence, and repetition rate provide an average power levelsubstantially similar to that of the commercial system, allowing for acomparison of the performance of the two systems.

Extracted human molar samples were used for ablation of enamel thereuponusing this experimental system. The molars were stored in householdbleach, and cleaned by hand with isopropyl alcohol. The molars weremounted on their sides, or by the root, in Plaster of Paris.Thermocouple holes were created with diamond tipped drills and sawsusing water soluble cutting fluid. Drilled molars were manually rinsedwith tap water to remove the cutting fluid. Minimal preparation of themolars was deliberate to mimic dental office applications.

The molar samples were irradiated with a Coherent Diamond™ 225i slablaser excited with a RF power supply model # D64/84 RF Amp/LC Filterfrom Coherent in Bloomfield, Conn. The laser contains optical beamcorrection components so the beam quality is specified at a M2 of 1.3.The laser beam was directed through and modulated by an AO switch, model# AGM-4010AJ1MD with RF power supply model # E41277, both fromIntraAction Corporation of Bellwood, Ill. The laser modulated pulseswere delivered through an articulating arm model # PLATA1042 from LaserMechanisms of Novi, Mich. The laser was focused to a spot size of 169 μmin diameter using a 78 mm FL plano/spherical lens. Some of the molarswere mounted in Plaster of Paris on their sides so the cutting trialswere performed on the relatively flat side of the molar eliminating thedifference of geometry of various molar cusp regions.

The molars were hydrated during cutting by a CoolMist Portamist Model #60M12 using compressed air at 135 psi. The mist angle was approximately45° to the cutting angle, the cut angle (i.e., the angle of incidence ofthe laser beam) being nearly 90° i.e., vertical or perpendicular to thecutting surface. It should be understood, however, that these angles areillustrative and cut angles other than 90° and mist angles other than45° are also within the scope of the invention.

The laser pulse rise time was measured to be 75 μsec, so that all laserpulsing was initiated for 75 μsec. Then the AO Switch was triggered tocreate 5 to 20 μsec pulse widths with no rise or fall time. The averagepower of the 225i laser was checked at the factory by Coherent, andconfirmed in the final configuration with both an Ophir model #FL-300A-LP-SH and model # F-150A-SH-V1-ROHS connected to an Ophir Nova,serial # 44797, optical power display meter. Pulse durations of 5, 10,15, and 20 μsec were used with repetition rates of 250, 500, 1000, and1500 Hertz for cutting the enamel of the molars. The laser andperformance parameters of this system, as measured during the operation,are presented in a table in FIG. 5.

Some of the extracted human molars were mounted with their roots mountedin Plaster of Paris. Teeth were cut on the molar cusps where thegeometry varies and the enamel thickness is the greatest. The molarswere hydrated as described above. The laser optical and beam deliveryarrangement described above was used, except a 20 μsec pulse length anda 2 kHz repetition rate was used.

Thermocouples were mounted at various locations in the molars to measurethe rise in temperature at various locations in the interior of themolars. One thermocouple was mounted up through the root of the molar.Three other thermocouples were mounted below, at, and above the enamelline respectively. Prefabricated insulated thermocouples, type KChromel-Alumel, 0.05″ diameter cable, with a miniature K-Type connectorwere used for temperature measurement. The thermocouples were mounted inthe drilled holes in good thermal contact and were held in place withthermally conductive epoxy, Resin Technology Group, Part # DP012209-1.The thermocouples voltages were converted to temperatures using a Fluke80TK Thermocouple Module, model # FLU890TK mounted directly into a Fluke77 Series digital voltmeter. The same optical arrangement as thatdescribed above was used, and 20 μsec pulse lengths at 2 kHz repetitionrates were utilized, cutting continuously for three minutes.

Two additional molars were mounted by the root in Plaster of Paris andcut with the same optical system and hydrated as described above. A 20μsec pulse length and 2 kHz repetition rate was used for the cutting ofthese two molars.

Sixteen molars were cut to measure the rate of removal of enamel byweight. FIG. 6 depicts the removal rate for the sixteen molars, whereeach molar is cut using a unique combination of pulse length andrepetition rate (e.g., 5 μsec and 250 Hz, 5 μsec and 1.5 kHz, 15 μsecand 250 Hz, 15 μsec and 1 kHz, 20 μsec and 1.5 kHz, etc.). FIG. 6 alsoshows the average rate of cutting by weight using a conventional system,such as the Erbium dental laser. The projected achievable cutting rates,obtained by linear interpolation are also shown in FIG. 6.

In addition to the laser beam parameters (e.g., pulse length, repetitionrate, pulse energy, fluence at focus, etc.) the variables that canimpact the enamel cutting speed include flatness of the side of themolars, the elapsed time since a molar was extracted, and the steadinessof the hand-held laser system cutting tip. When a pulse repetition rateof about 1 kHz or greater is used, scanning of the targeted region ofthe hard tissue may be used. Substantially steadily directing a laserbeam to the targeted region can impact the cutting speed. In general,however, this experimental system was about five times faster than theknown systems including the Er:YAG and Er:YSGG systems that havevolumetric removal rates of approximately 0.31 and 0.33 cm³/min.

Five molars were cut to measure the rate of removal of enamel by volume.These molars were irradiated with a laser beam configured to deliverspulses of 20 μsec pulse length, 2 kHz repetition rate, 9.7 mJ pulseenergy, and 21 J/cm² fluence at focus. The typical average density ofenamel, e.g., carbonated hydroxyapatite, is 2.94 g/cm³. The removalrates (by weight) are listed in the table presented in FIG. 7.

The spot size (i.e., area of the treatment region) and energy profileover the spot size are also important parameters. For example, even at alower fluence, higher cutting rates were achieved at a spot size ofabout 250 μm in diameter, which may yield a fluence of about 10 J/cm²compared to a smaller spot size (e.g., 169 μm in diameter), yielding afluence of about 20 J/cm². One of the reasons for the increased cuttingspeed at the lower fluence is the energy distribution across the spotsize. When a plano/convex focus lens is used, the focused spot energyprofile may be Gaussian i.e., a substantial amount of pulse energy isdelivered at or near the focal point of the lens, which is usually anarea near the center of the spot. A comparatively small amount of energyis delivered to the remaining area of the spot. In some instances, theenergy delivered to the focal region of the spot can be above theablation threshold while that delivered to the non-focal region may bebelow the ablation threshold. The energy in the “wings of the Gaussianenergy pulse profile” (i.e., the energy directed to the non-focalregions) which is below the ablation energy can be absorbed as residualenergy causing undesired heating of the tissue.

Thus, cutting when the enamel was positioned substantially at the focuscan cause small craters below the cut surface. By positioning the enamelslightly above the focus, the amount of energy delivered at or near thefocal region can be reduced, and, accordingly, less energy is consumedin ablating portions of the enamel below the cut surface.Correspondingly more pulse energy can be available for ablation in thenon-focal region of the spot size, thereby achieving substantiallyuniform ablation in the target spot. Aspheric adaptive beam shapingoptics may also be used to create a “flat top” or “top hat” energydistribution at focus, such that relatively less pulse energy isdelivered near the focal region and relatively more energy may bedelivered to the non-focal region of the spot, so that substantially allpulse energy is utilized for ablating the surface of the hard tissue.Alternatively or in addition, adaptive optics may be used to modify theoptical pulse energy distribution to be non-Gaussian, substantiallymaintaining the energy in the entire spot above ablation threshold,thereby reducing the amount of residual energy consumed.

The performance of this system according to the present invention wascompared with that of the commercial system. Charing occurred in thetooth when the commercial system was used. Even though this exemplarysystem cut tooth enamel at about five times faster than the commercialsystem, cutting using this exemplary system did not cause anysubstantial increase in temperature of adjacent tissue, and no charring.Moreover, due to the larger pulse lengths and larger pulse energy, thecommercial system was less precise compared to the exemplary system. Thelarger pulse energies of the commercial system may also have affectedthe texture of the tooth. For example, at the bottom of a tooth cutusing the commercial system, a section of enamel separated and liftedoff the dentin. The separation may have been caused by the undesiredheating of the enamel which thermally expands faster than the dentin andmay mechanically separate therefrom. Accordingly, the instant inventionprovides higher removal rates coupled with more precise removallocation, yielding a much more suitable dental treatment system thanpresently available.

In order to measure the heating of the internal tissue of the molarscaused by this exemplary system, four molars were cut with thermocouplesmounted therein as described above. In a first molar the thermocouplewas located in the nerve chamber. In a second molar the thermocouple waslocated below the enamel line. In a third molar the thermocouple waslocated at the enamel line. In a fourth molar the thermocouple waslocated above the enamel.

The temperature profile of the four molars is shown in FIG. 8. In athermocouple mounted approximately 2 mm above the nerve chamber, at thebeginning of the treatment, the temperature of the molar's nerve chamberdropped due to the convection effect of the water mist on the molar andthen rose slightly only reaching substantially the initial temperature.In vivo molars are at higher temperatures and the water mist will haveeven a more significant cooling effect on these molars. In general, itis desirable that the procedure not increase the temperature ofremaining material substantially. For example, a local temperature riseof up to about 5° C. may be suitable, as an upper limit for the pulptemperature rise. Lower temperature rise values are generally preferred,to provide patient comfort (e.g., up to about 1-3° C.).

The cutting experiments as described above show that the exemplary lasersystem can cut at least as quickly as a conventional rotating dentalhand-piece burr and much more quickly than conventional dental lasersystems. The cutting rates corresponding to this exemplary systemaccording to the invention, a conventional burr drilling system, and aconventional dental laser system are shown in FIG. 9.

While the foregoing has described what are considered to be the bestmode and/or other preferred embodiments, it is understood that variousmodifications can be made therein and that the invention or inventionsmay be implemented in various forms and embodiments, and that they maybe applied in numerous applications, only some of which have beendescribed herein. For example, it is possible that the described laserand laser system could be used in other (non-dental) applications, suchas cutting or contouring bone. It is intended by the following claims toclaim that which is literally described and all equivalents thereto,including all modifications and variations that fall within the scope ofeach claim.

1. A method for treating hard tissue, the method comprising the stepsof: generating a laser beam with a repetition rate greater than about0.5 kHz; directing the laser beam at an exposed surface of the hardtissue; and removing at least a portion of the exposed surface of thehard tissue without substantially increasing temperature of adjacenttissue.
 2. The method of claim 1, wherein the laser beam comprises awavelength in a range of about 9 μm to about 10 μm.
 3. The method ofclaim 1, wherein the laser beam comprises a pulse length in a range ofup to about 30 μsec.
 4. The method of claim 1, wherein the laser beamcomprises a fluence up to about 50 J/cm².
 5. The method of claim 1,wherein the laser beam comprises a repetition rate up to about 4 kHz. 6.The method of claim 1, wherein the laser beam has a nominal diameter ofup to about 2000 μm.
 7. The method of claim 1, wherein the laser beam isgenerated by a CO₂ laser.
 8. The method of claim 7, wherein the CO₂laser operates with a gas in a range of about 260 Torr to about 600Torr.
 9. The method of claim 8, wherein the gas is selected from thegroup consisting of ¹²C(¹⁸O)₂ gas and ¹²C(¹⁶O)₂ gas.
 10. The method ofclaim 1, wherein the laser beam is directed at the exposed surface ofthe hard tissue utilizing a hollow waveguide.
 11. The method of claim10, wherein positioning of the hollow waveguide relative to the exposedsurface of the hard tissue is manually controlled.
 12. The method ofclaim 10, wherein positioning of the hollow waveguide relative to theexposed surface of the hard tissue is robotically controlled.
 13. Themethod of claim 1, wherein the laser beam is moved relative to theexposed surface of the hard tissue to remove incrementally the portionof the exposed surface of the hard tissue.
 14. The method of claim 1,wherein a remaining portion of the hard tissue does not exhibit anysubstantial charring.
 15. The method of claim 1, wherein the hard tissueis removed at a rate of up to about 0.5 g/sec.
 16. The method of claim1, wherein the hard tissue is removed at a volumetric rate of up toabout 1.7 cm³/sec.
 17. The method of claim 1, wherein the laser beamcomprises a pulse energy in a range of up to about 30 mJ.
 18. The methodof claim 1, wherein the hard tissue is removed by explosivevaporization.
 19. The method of claim 1, further comprising the step ofdirecting a fluidic flow at the exposed surface of the hard tissue. 20.The method of claim 19, wherein the fluidic flow is selected from thegroup consisting of a fluid, a gas, and combinations thereof.
 21. Themethod of claim 19, wherein the fluidic flow is adapted to at least oneof: contribute to controlling temperature of the hard tissue; andcontribute to reducing redeposition of removed material onto the exposedsurface of the hard tissue.
 22. The method of claim 19, wherein thefluidic flow is directed to the exposed surface of the hard tissue at anangle of incidence in a range from about zero degrees to about 90degrees.
 23. The method of claim 1, wherein the hard tissue us selectedfrom the group consisting of dental enamel, dentin, and bone.
 24. Themethod of claim 1, wherein the exposed surface of the hard tissue atwhich the laser beam is directed during a single pulse has across-sectional area of up to about 0.03 cm².
 25. The method of claim 1,wherein the step of directing the laser beam comprises delivering asequences of pulses for up to about three minutes.
 26. The method ofclaim 1, wherein the laser beam is directed to the exposed surface ofthe hard tissue at an angle of incidence up to about 90 degrees.
 27. Themethod of claim 1, wherein the step of directing the laser beamcomprises: focusing the laser beam using a lens having a focal length;and positioning the exposed surface of the hard tissue relative to thelens at a distance slightly less than the focal length.
 28. The methodof claim 1, wherein the step of directing the laser beam comprisesmodifying the laser beam using an optical system, such that a profile ofthe modified laser beam is non Gaussian.
 29. A system for treating hardtissue, the system comprising: a laser beam generator for generating alaser beam with a repetition rate greater than about 0.5 kHz; an opticalcomponent for directing the laser beam at an exposed surface of the hardtissue; and a laser beam controller for controlling at least one laserbeam parameter, such that at least a portion of the exposed surface ofthe hard tissue is removed without substantially increasing atemperature of adjacent tissue.
 30. The system of claim 29, wherein thelaser beam parameter is a wavelength in a range of about 9 μm to about10 μm.
 31. The system of claim 29, wherein the laser beam parameter is apulse length in a range of up to about 30 μsec.
 32. The system of claim29, wherein the laser beam parameter a fluence up to about 50 J/cm². 33.The system of claim 29, wherein the laser beam parameter a repetitionrate up to about 4 kHz.
 34. The system of claim 29, wherein the laserbeam has a nominal diameter of up to about 2000 μm.
 35. The system ofclaim 29, wherein the laser beam generator is a CO₂ laser beamgenerator.
 36. The system of claim 35, wherein the CO₂ laser beamgenerator operates with a gas in a range of about 260 Torr to about 600Torr.
 37. The system of claim 36, wherein the gas is selected from thegroup consisting of ¹² _(C)(¹⁸O)₂ gas and ¹²C(¹⁶O)₂ gas.
 38. The systemof claim 29, wherein the optical component comprises a hollow waveguide.39. The system of claim 38, further comprising a handpiece for manuallypositioning the hollow waveguide relative to the exposed surface of thehard tissue.
 40. The system of claim 38, further comprising a roboticcontroller for positioning the hollow waveguide relative to the exposedsurface of the hard tissue.
 41. The system of claim 40, wherein therobotic controller is configured to move the laser beam relative to theexposed surface of the hard tissue to remove incrementally the portionof the exposed surface of the hard tissue.
 42. The system of claim 29,wherein a remaining portion of the hard tissue does not exhibit anysubstantial charring.
 43. The system of claim 29, wherein the hardtissue is removed at a rate of up to about 0.5 g/sec.
 44. The system ofclaim 29, wherein the hard tissue is removed at a volumetric rate of upto about 1.7 cm³/sec.
 45. The system of claim 29, wherein the laser beamcomprises a pulse energy in a range of up to about 30 mJ.
 46. The systemof claim 29, wherein the hard tissue is removed by explosivevaporization.
 47. The system of claim 29, further comprising a fluiddispenser for directing a fluidic flow at the exposed surface of thehard tissue.
 48. The system of claim 47, wherein the fluidic flowdispensed by the fluid dispenser is selected from the group consistingof a fluid, a gas, and combinations thereof.
 49. The system of claim 47,wherein the fluidic flow is adapted to at least one of: contribute tocontrolling temperature of the hard tissue; and contribute to reducingredeposition of removed material onto the exposed surface of the hardtissue.
 50. The system of claim 47, wherein the fluid dispenser directsthe fluidic flow to the exposed surface of the hard tissue at an angleof incidence in a range from about zero degrees to about 90 degrees. 51.The system of claim 29, wherein the hard tissue is selected from thegroup consisting of dental enamel, dentin, and bone.
 52. The system ofclaim 29, wherein the exposed surface of the hard tissue at which thelaser beam is directed during a single pulse has a cross-sectional areaof up to about 0.03 cm².
 53. The system of claim 29, wherein the laserbeam controller configures the laser beam as a sequences of pulses forup to about three minutes.
 54. The system of claim 29, wherein theoptical component is configured to direct the laser beam to the exposedsurface of the hard tissue at an angle of incidence up to about 90degrees.
 55. The system of claim 29, wherein the optical componentcomprises a lens having a focal length for focusing the laser beam, suchthat the exposed surface of the hard tissue relative to the lens is at adistance slightly less than the focal length.
 56. The system of claim29, wherein the optical component modifies the laser beam, such that aprofile of the modified laser beam is non Gaussian.